Sensor employing impedance measurements

ABSTRACT

A sensor uses an immobilized affinity component capable of interacting with analyte species and being associated with a conducting polymer such that the interaction of the affinity component and the analyte induces change in the electrical properties of the polymer. An AC signal is applied to the polymer, and the induced change in impedance resulting from the interaction is measured. The impedance is measured at a frequency or frequencies corresponding to a peak or peaks in the relationship between frequency and impedance change for the polymer and the analyte. The measurement may be made by reference to the imaginary or real component of impedance. The polymer may be in the form of a layer bridging two electrodes between which the impedance is measured. The two electrodes may together define an interdigitated electrode assembly.

BACKGROUND OF THE INVENTION

1. Field of the Invention

The present invention relates to a sensor and more particularly but notexclusively, to a biosensor.

2. Related Art

Biosensors are used for determining the presence and/or amount of ananalyte in a sample. Typically a biosensor comprises a bio-component(e.g. enzyme) which is specific for the analyte to be determined andwhich interacts therewith to produce a detectable change indicative ofthe presence and/or amount of the analyte in a sample. A particularexample of biosensor comprises an enzyme electrode system whereby thecatalytic reaction of suitable redox enzyme on analyte produces anelectroactive species which is detected electrochemically. A glucosesensor using glucose oxidase as the bio-component is a specific exampleof an enzyme electrode system. Here glucose is oxidised by oxygen viaglucose oxidase to produce gluconic acid and hydrogen peroxide. thelatter being detected electrochemically. Systems based on oxidaseenzymes are particularly well suited to biosensors because the oxygenacts as the mediator (i.e. to shuttle electrons to the electrode fromthe enzyme active site) and no additional reagents are required (i.e. itis a reagentless system). However, not all analytes have correspondingoxidase enzymes therefore limiting the number of substrates that can bedetermined using this system. Dehydrogenase enzymes are more diversethan oxidases and hence enable that a wider variety of analytes can bemeasured by biosensor systems. A key disadvantage of dehydrogenaseenzymes, however, is that oxygen cannot act as a mediating (electronaccepting) species and instead soluble nicatinamide adenine dinucleotide(NAD(P)⁺/NAD(P)H) cofactors are required to act as electronacceptors/donors. Thus, biosensor systems based on dehydrogenase enzymesare not reagentless which restricts practical application.

It is known from European Patent Specification No. 402917 to provide abiosensor employing a thin surfactant polymeric electrically conductinglayer to which may be bound members of specific binding pairs. Bindingof an analyte to the specific binding pair member layer changes theelectrical properties of the layer to enable the detection of theanalyte. A DC voltage is applied between a pair of electrodes bridged bythe layer, and changes in the voltage are measured. Alternatively,electrical alternating current measurements may be used for filteringout background noise due to non-specific matrix effects. Unfortunately,the sensitivity of such biosensors is limited as the measured changes inthe electrical properties of the layer are relatively small.

SUMMARY OF THE INVENTION

The object of the present invention is to obviate or mitigate theabovementioned disadvantage.

According to the present invention there is provided a sensor comprisingan immobilised affinity component capable of interacting with an analytespecies and being associated with a conducting polymer such that theinteraction of the affinity component and the analyte induces a changein the electrical properties of the polymer, means for applying an ACsignal to the polymer, and means for detecting the response of thepolymer to the applied signal to detect the induced change, wherein thedetecting means comprises means for measuring the impedance of thepolymer at a frequency corresponding to a peak in the relationshipbetween frequency and impedance change for the polymer and the analyte.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 depicts the activity of an immobolised enzyme versus time for afirst exemplary sensor embodiment;

FIG. 2 depicts an impedance change versus applied voltage response topyruvate for the first exemplary sensor;

FIG. 3 depicts impedance versus time for the first exemplary sensor inresponse to pyruvate and then NADH;

FIG. 4 depicts the exemplary first sensors resistance versus pH;

FIG. 5 depicts the exemplary first sensor's current versus voltageresponse in the presence and absence of pyruvate;

FIG. 6 depicts the impedance change versus time for a second exemplarysensor embodiment in response to ATP; and

FIGS. 7 and 8 respectively depict the real and imaginary components ofimpedance versus frequency of applied voltage for a third exemplarysensor embodiment.

DETAILED DESCRIPTION OF EXEMPLARY EMBODIMENTS.

The present invention is based on the realisation that impedance changesare very frequency dependent and therefore sensitivity may be greatlyenhanced by measuring the impedance to AC signals at a frequency orfrequencies at which the change in impedance resulting from exposure tothe analyte peaks.

The sensor of the present invention employs a conducting polymer toamplify, and transduce, the interaction between analyte and immobilisedaffinity component. During analyte interaction with the affinitycomponent the electrochemical properties of the conducting polymer arealtered and the changes detected, by measurement of impedance changes.Thus the interaction of the analyte species is directly transduced intoa measurable signal with no need for formation of an electroactiveproduct as in the case of conventional enzyme electrodes. The identityof the affinity component and the analyte species of interest and thestructure of the conducting polymers, will determine the nature of thechange in electrical properties which will result from the presence ofthe analyte. If the change is for example primarily capacitive, theimaginary component of the measured impedance is likely to exhibit apeak at a predetermined frequency and accordingly the imaginarycomponent of the impedance should be monitored at that predeterminedfrequency. Conversely, if the change is primarily resistive, it may bemore appropriate to monitor the real component of the impedance.

Conducting polymers are organic semi-conducting materials which can bereversibly switched through conducting and insulating states by varyingthe degree of oxidation, protonation and polymer chain orientation. Inthe sensor of the invention it is the interaction of the analyte withthe affinity component which. we believe, alters the local electrostaticinteractions between polymer and immobilised affinity component inducingchanges in polymer orientation (hence electrical properties i.e. chargemobility).

The affinity component will preferably be a macromolecule.

Preferably the sensor of the invention is a biosensor so that theaffinity component is a biocomponent. for example an enzyme, such askinase or dehydrogenase, or L(+) lactic dehydrogenase suitable formeasuring pyruvate. Alternatively the affinity component may be anantibody, antigen, lectin or receptor.

The polymer may be in the form of a layer bridging two electrodesbetween which the impedance is measured. The two electrodes may togetherdefine an interdigitated electrode assembly. The polymer layer may havea thickness of less than 5 μm, and may contain or incorporate a layerof, immobilised binding agent. The affinity component may be entrappedwithin a gel layer provided oil the conducting polymer layer on the sidethereof remote from the electrodes, the gel thickness being less than 50μm. Alternatively, the affinity component may be immobilised within theconducting polymer.

The conducting polymer may be poly(3-methylthiophene) or poly(pyrrole).

Thus, in one embodiment the affinity molecule is an enzyme wherebysubtle changes in electrostatic potential of the protein structureduring analyte interaction induces changes in the local environment incontact with the polymer film which may be detected through measuringthe impedance of the supporting conducting polymer.

Enzymes which can be used as the affinity component arc those thatrequire two or more components to achieve enzyme catalytic activity.That is to observe the interaction of analyte with enzyme no catalyticactivity should occur and the enzyme should merely act as an affinitysite (i.e. analogous to antibody and antigen interaction). For example.dehydrogenase enzymes arc a specific enzyme class that may find utilityin the described sensors whereby in the absence of solubleNAD(P)⁺/NAD(P)H cofactor the interaction of substrate (analyte) with theaffinity site of the enzyme is measured. Dehydrogenases for lactate,pyruvate, citrate and malate are possible examples of enzymes which maybe used. A further class of enzyme are kinases where both the analyteand phosphate donating ATP are required to permit catalytic activity. Anexample is with hexokiniase which catalyses the phosphorylation ofglucose by ATP. In the absence of glucose the interaction of ATP withhexokinase immobilised over a conducting polymer film induces a changein polymer conductivity. The use of other enzymes such as oxidaseslipases, hydratases and proteinases may be restricted as these systemsdo not require the addition of dual substrates. For example withoxidases the addition of analyte will furness the catalytic activity ofthe enzyme unless oxygen is totally excluded. In such circumstances theenzyme is acting catalytically and not as an affinity site which leadsto exclusion in the described sensor format. Oxidases, however may beused if the FAD prosthetic group (responsible for electron transfer fromsubstrate to oxygen) is removed (apo-enzyme). Under such circumstancesthe enzyme now acts as an affinity site for substrate due to itsinability to interact with molecular oxygen.

If the affinity molecules are antibodies, antigens lectins and receptorsrather than enzymes, the interaction between analyte and affinity sitescan be detected directly unlike the multi-step protocols requiringindicator labels (e.g. radioisotope, enzyme conjugate) currently used instandard analysis. Antibodies can be used to detect small molecules suchas pesticides (e.g,. dinitropentol derivatives), antiepipeptic(phenytoin) and restricted drugs (e.g. morphine, heroin). Alternatively,the target may be protein (albumin) or hormone (leutenisinig hormone).Lectin substrates can be sugar residues, for example, Concanavilin A forglucose/dextrin binding.

Antigens may also be immobilised to act is the affinity component in thedetection of the corresponding antibody or receptor. For example,leutenising hormone could be immobilised and used to detectanti-leutenising hormone in solution. Low molecular weight antigens(e.g. drugs) may be used as affinity molecules provided that covalentlinkage of the molecule to the surface of the conducting polymer isundertaken. This will provide accessibility to allow anitibody/receptorinteraction and in addition assure retention of the affinity molecule.

Studies so far have indicated that the conducting polymer type has aninfluence over detection of analyte interaction with immobilisedaffinity molecule. For example work using L(+) lactic dehydrogenase hasindicated that poly(3-methylthiophene) (as opposed to poly(pyrrole) andpoly(aniline)) is suitable for detecting interaction of pyruvate analytewith immobilised L(+) lactic dehydrogenase. In a further example,interaction of leuteinising hormone with immobilised antibody(anti-leuteinising hormone) can be detected when poly(pyrrole) is thesupporting conducting polymer.

The reasons for the apparent different sensitivities of conductingpolymer films are at this time unclear. However, it must be noted thatthe properties of conducting polymers do differ. For example,poly(3-methylthiophene) has a more ordered and flexible structurecompared to that of poly(pyrrole) which could be a factor in determiningthe sensitivity of the sensor. The nature of the affinity molecule andanalyte may also contribute to the sensor signal. That is the physicalsize of the analyte and the magnitude of the induced changes in theaffinity molecule will all contribute to the observed response.

Preferably the conducting polymer used should permit immobilisation ofthe affinity molecule during the electropolymerisation process. Thisprovides intermittent contact between the polymer film and bio-affinitymolecule which may influence the magnitude of the response. Thuspoly(pyrrole) is preferred because electropolymerisation can be achievedunder neutral aqueous conditions thus permitting the entrapment of theaffinity molecule during polymer deposition. The polymerisation of3-methylthiophene, however, has to be performed in acidic non-aqueousconditions which compromises bio-affinity molecule stability. Thus, theimmobilisation of the bio-affinity molecule may subsequently be effectedby gel entrapment over the pre-formed conducting polymer film.Therefore, it is believed that the type of conducting polymer used is acompromise between analyte access to the bio-affinity molecule, theintimacy between the conducting polymer film and the bio-affinitymolecule, and the properties of the supporting conducting polymer toreport analyte interaction.

Therefore the type of conducting polymer used may be expected to dependon the type of bio-affinity molecule/analyte being monitored.

The interrogation technique used is impedance spectroscopy. Measurementsmade over a wide frequency range (typically 100 kHz-0.01 Hz) make itpossible to characterise different interfacial and bulk regions ofconducting polymer films. Impedance data can be represented in a numberof different forms all of which place a different weighting on parts ofthe frequency spectrum. Thus, by using an impedance function orparameter it is possible to highlight/amplify changes in conductingpolymer behaviour during analyte interaction with immobilised affinitymolecule. This is not possible with techniques such as amperometry wherethe recorded response is a summation of all the electrochemicalproperties (i.e. the processes cannot be separated). Surface potential(charge) has been measured, but this is probably not a viable selectiveway of identifying interaction between the analyte and the affinitycomponent, and only charged molecules can be assessed in this way.

Immobilisation of the affinity molecule is preferably achieved byentrapment within the conducting polymer film during theelectropolymerisation process. When this is not permitted due to theconditions required for polymerisation (i.e. acidic organic phase), thenthe affinity molecule may be adsorbed onto pre-formed conducting polymerfilm. The adsorbed affinity molecule may then be “fixed” by anoverlaying gel-like polymer, notably a hydrogel which allows theinclusion of water for analyte transport and subsequent interaction withthe immonbilised affinity molecule. A typical hydrogel would be poly(acrylamide). Covalent coupling of affinity molecule surface lysylresidues to conducting polymer derivatives is also not precluded as animmobilisation technique.

A sensor in accordance with the invention most preferably comprises aninterdigitated electrode whereby the conducting polymer layer or filmbridges the gap separating the electrode digits. The affinity moleculemay be either entrapped within the bridging film or alternativelyimmobilised on the side of the conducting polymer layer remote from theelectrode.

Conducting polymer thickness should be sufficient to bridge the gapbetween interdigitated electrodes (10 μm). The effect of film thicknesson tie responsiveness of the polymer layer has yet to be conclusivelydetermined, however, conducting polymer films thicker than 5 μm werefound to be unstable and readily detached from the interdigitatedelectrode surface.

To ensure rapid transfer of solute through the gel layer, gel thicknessis preferably less than 50 μm, but up to 200 μm is possible. The keyconsideration is adequate affinity component in contact with theconducting polymer.

The conducting polymer may conveniently be produced byelectropolymerisation.

The invention is illustrated by the following non-limiting Examples andFIGS. 1 to 8 of the accompanying drawings which illustrate results ofthe Examples.

EXAMPLE 1

A sensor was produced (using the procedure below) based on anInterdigitated gold electrode (IDE) (7 mm×7 mm including electrodecontact pads, 10 μm thick digits, 15 μm separation, total of 50 digitswith 25 connected to one pad and 25 connected to the other pad) andincorporating an immobilised layer of L(+)-lactic dehydrogenase (LDH) torepresent a model system. LDH catalyses the reversible reduction ofpyruvate with NADH cofactor acting as the electron donor.

Poly(3-methylthiophene) was deposited onto the IDE viaelectropolymerisation from a degassed acetonitrile solution containing0.1M HClO₄ and 0.1M 3-methylthiophene. Electropolymerisation wasachieved through potential scanning between −0.4-1.5V (vs SCE) at a scanrate of 100 mV/s.

LDH (250 Units) was adsorbed onto the surface of thepoly(3-methylthiophene) coated IDE. 5 μl acrylamide/bis-acrylamide (10%w/v) containing 1 μl TMED was then applied to the surface of the coatedIDE. Polymerisation of acrylamide/bis-acrylamide was achieved throughthe addition of 5 μl ammonium persulphate (10% w/v). The polymerisationprocess is complete within 2 min and results in a strongly adherentfilm.

The resultant sensor was then subjected to a number of tests.

In order to demonstrate that the LDH had not been denatured during theimmobilisation process UV spectroscopy was used to measure the enzymeactivity of the sensor. The assay consisted of phosphate buffer (pH 6,50 mM) containing 1 mM NADH. The IDE was submerged in the reaction mixand the reaction initiated by the addition of 5 mM sodium pyruvate. NADHoxidation was recorded at 340 nm. A control experiment was conductedwithout the electrode present in the reaction mix. The results are shownin FIG. 1 which clearly demonstrates that the immobilised enzymemaintained its activity as witnessed by the reduction in absorbance byNADH with time.

Various impedance measurements were performed using a Voltech frequencyresponse analyser connected to a PARC 273 potentiostat. Measurementswere performed in phosphate buffer (pH 6, 50 mM containing 50 mM KCl)previously deoxygenated by N₂. The impedance of the enzyme IDE electrodewas measured from a starting frequency of 1 Hz and ended at 100,000 Hzat an amplitude of 20 mV. Raw data obtained from the frequency responseanalyser were presented as a output voltage:input voltage ratio whichwas used to calculate the real part (Z_(r)) and imaginary part (Z_(im))of the impedance from which the total impedance (IZI) was derived.

Exposure of the modified IDE to pyruvate caused a decrease in the ratio(thus impedance) indicating an increase in conductivity particularly inthe low frequency region. A plot of Z_(im) vs frequency of the same dataresulted in a dispersion/relaxation (i.e. peak) at a particularfrequency and that frequency was used to extract the impedance. Thatfrequency was chosen because dispersions/relaxations are indicative ofspecific processes occurring within the conducting polymer.

FIG. 2 illustrates the effect of bias potential on the response of thesensol to pyruvate. Measurements were performed in saline phosphatebuffer (pH 6, 50 mM containing 50 mM KCl). The impedance of the modifiedIDE was allowed to attain steady state at each potential and theresponse to 1 mM sodium pyruvate recorded. At potentials more positivethan −0.4 V (vs SCE) a decrease in polymer impedance (resistance) wasobserved (FIG. 2) but below at more negative potentials the polymerbecame less conductive following the addition of pyruvate. Use of −0.4V(vs SCE), however, gave the best resolution of impedance plots. Amodified IDE containing no enzyme showed relatively negligible impedancechanges (FIG. 2) across the same potential range. Thus FIG. 2 shows thatthe conductivity of the polymer can be modulated by altering the appliedbias potential.

The magnitude of the impedance change was also affected by the pH of thebathing solution with an optimum response being observed at pH 5.5.

FIG. 3 illustrates the dynamic response of a modified IDE to pyruvate (1mM) and NADH (0.25 mM). Following the addition of pyruvate a rapid dropin polymer impedance was found which attained a steady state value (FIG.3). This suggests that pyruvate binding to immobilised LDH is anequilibrium process. The addition of NADH co-factor caused a furtherimpedance drop which could be due to polymer oxidation via NAD⁺ (formedduring the course of the enzyme reaction).

FIG. 4 illustrates the dependence of resistance change on pH for apoly(met)/LDH interdigitated electrode.

Square wave voltammograms (FIG. 5) of the sensor in the presence (A) andabsence of (B) of pyruvate (1 mM) in saline phosphate bufferdemonstrates that the impedance change was a non-Faradaic process (i.e.there was no increase in the peak current recorded in the presence ofpyruvate). This removes the possibility of direct electron transfer fromthe redox centre of LDH to the supporting conducting polymer. In theabsence of a Faradaic reaction it is plausible to suggest that theobserved changes in polymer impedance during the interaction of pyruvatewith immobilised LDH may be due to re-orientation of the conductingpolymer chains. It still remains unclear as to how enzyme immobilised onthe polymer surface can affect the bulk conducting polymer properties.

The type of immobilisation gel used has had an effect on the pyruvateresponse of the sensor. When polyacrylamide was replaced with thethermoset gel, poly(vinyl alcohol), an impedance increase was observed(as opposed to an impedance decrease as found when usingpoly(arylamide). The reason for this significant alteration in responseremains unclear but it is most probably that the interaction with theconducting polymer film will be different for poly(acrylamide) andpoly(vinyl alcohol). That is to say the interaction of gel phase withthe conducting polymer influences the nature of the interaction with thebio-affinity molecule. A problem associated with the use of poly(vinylalcohol) as the gel phase is one of stability whereby the gel layerdetaches from the IDE coated electrode after a prolonged exposure inaqueous solution. In this respect poly(acrylamide) is the preferredchoice of hydrogel.

EXAMPLE 2

A further example of a class of enzyme that can find utility in thedescribed sensor is given by work performed using hexokinase. Hexokinasecatalyses the phosphorylation of glucose by ATP. The same sensor formatwas used as previously described save in this case 250 units ofhexokinase replaced LDH. FIG. 6 illustrates the response to ATP of apoly(3-methylthiophene) coated electrode containing immobilisedhexokinase. It can be seen that the addition of ATP resulted in animpedance (resistance) increase in (as opposed to a decrease in the caseof pyruvate/LDH interaction). Saturation of the enzyme sites appears tooccur at very low ATP (below 0.05 mM) concentrations which could be dueto the affinity constants of the enzyme to ATP substrate. The differentresponse between the pyruvate/LDH and ATP/hexokinase could be attributedto several factors. For example, different enzymes may undergo alternatere-orientation effects.

The LDH and hexokinase model systems demonstrate that the sensor formatdescribed may be used to detect further analytes with use of theappropriate enzyme.

EXAMPLE 3

A reagentless immunosensor for leutenising hormone (fertility hormone)has been demonstrated using a similar sensor format. That is measuringthe impedance changes of a conducting polymer (containing antibody toleutenising hormone) bridging an IDE. The important difference in thisexample is that it was possible to record a response using antibodyentrapped in poly(pyrrole). By using poly(pyrrole) the entrapment ofantibody during polymer deposition was possible by virtue that thepolymerisation process could be performed under neutral aqueousconditions which is compatible to antibody stability. In the case of theleutenising hormone sensor the overlaying gel phase would not beexpected to restrict diffusion of the leutenising hormone antigen to theimmobilised antibody.

Upon exposure of the sensor to leutenising hormone antigen changes inthe impedance properties of the supporting poly(pyrrole) polymer wererecorded. Those changes are represented in FIGS. 7 and 8. FIG. 7 showingthe real component of the impedance across a range of frequencies beforeexposure (t=0) and 15 minutes after exposure (t=15), and FIG. 8 showingthe imaginary component in the same circumstances. It will be seen thatthere is a pronounced peak in the imaginary component in the frequencyrange of from 10 Hz to 100 Hz before exposure, the peak disappearingafter exposure. Accordingly measurements will be made at a frequency orfrequencies within this range to maximise the sensitivity of the system.The exact frequency selected may be selected to maximise the change thatoccurs between the measurements before and after exposure.

Thus FIG. 8 shows that changes in the polymer will be highlighted bychanges in the magnitude and position of the peak (which may be referredto as the relaxation point). Although larger responses can be obtainedat frequencies to either side of this relaxation peak those responsesresult at least in part from processes not dependent upon interactionswith the affinity component associated with the conducting polymer. Theoverall sensor structure can be considered as being made up of a seriesof layers which exhibit different electrical characteristics. A firstlayer corresponds to the solution:polymer interface and the impedance ofthat first layer is a function of the mobility of ions moving in and outof the polymer. This process is typically capacitive in nature. A secondlayer is the polymer itself, the electrical characteristics of which areof prime interest. A third layer is the electrode:polymer interface theimpedance characteristics of which depend upon processes including thetransfer of electrons between the polymer and the electrode. This is apredominantly resistive process. Using an interdigitated electrodeformat minimises the polymer:solution interface contribution. In theillustrated example, processes taking place in the polymer can beidentified by reference to a middle frequency range. Looking at a lowerfrequency would detect substantial contributions from changes in thesolution resulting for example from the introduction of analyte. Itshould be pointed out however that an even more extended frequency rangethan that shown in FIGS. 7 and 8 could be applied to identify otherpeaks relating to polymer processes.

It will be noted that the real component of impedance in the examplerepresented in FIG. 7 does not exhibit any frequency-dependent peaks.Accordingly in this example it is sensible to concentrate attention ofthe imaginary component of the impedance. In other examples howeverpeaks may occur in the real component of the impedance, and in suchexamples measurements will be made of the real rather than the imaginarycomponent of the impedance.

Sensors in accordance with the present invention are in effect “tuned”to exhibit high sensitivity given the nature of the polymer, theaffinity component and the analyte. A sensor may be pre-tuned to exhibitmaximum sensitivity to a particular analyte such that a signal isapplied to the polymer at a fixed and predetermined frequency andmeasurements are always made at that frequency. Alternatively, a sensorin accordance with the invention could be “self-tuning”, that issoftware would be provided to apply a wide range of frequencies to thepolymer, to record changes in the real and imaginary components of theimpedance over time after the sensor is exposed to the analyte, and toanalyse data recorded at frequencies corresponding to peaks in therelationship between frequency and measured impedance. The resultantdata could then be compared with a database of information representingthe response of the sensor to exposure to known analytes with a view toidentifying an unknown analyte.

What is claimed is:
 1. A sensor comprising: an immobilised affinitycomponent capable of interacting with a predetermined analyte speciesand being associated with a conducting polymer such that interaction ofthe affinity component and the analyte induces a change in electricalproperties of the polymer, means for applying an AC signal to thepolymer, and means for detecting the response of the polymer to theapplied signal to detect the induced change, wherein the detecting meanscomprises means for measuring the impedance of the polymer at afrequency selected to correspond to a peak in the relationship betweenfrequency and impedance change for the polymer and the predeterminedanalyte.
 2. A sensor as in claim 1, wherein the imaginary component ofimpedance is measured at a frequency at which there is a peak in therelationship between the imaginary component and frequency.
 3. A sensoras in claim 1 wherein the real component of impedance is measured at afrequency at which there is a peak in the relationship between the realcomponent and frequency.
 4. A sensor as in claim 1 wherein the polymeris in the form of a layer bridging two electrodes between which theimpedance is measured.
 5. A sensor as in claim 4, wherein the twoelectrodes together define an interdigitated electrode assembly.
 6. Asensor as in claim 4, wherein the polymer layer has a thickness of lessthan 5 μm.
 7. A sensor as in claim 4, wherein the layer of theconducting polymer contains, or incorporates a layer of, immobilisedaffinity component.
 8. A sensor as in claim 7, wherein the affinitycomponent is entrapped within a gel layer provided on the conductingpolymer layer on the side thereof remote from the electrodes and the gelthickness is less than 50 μm.
 9. A sensor as in claim 1 wherein theaffinity component is immobilised within the conducting polymer.
 10. Asensor as in claim 1, wherein the affinity component is immobilised byentrapment in a gel provided on the conducting polymer.
 11. A sensor asin claim 1, wherein the conducting polymer is poly(3-methylthiophene).12. A sensor as in claim 1 wherein the affinity component is amacromolecule.
 13. A sensor as in claim 1 wherein the affinity componentis a biocomponent.
 14. A sensor as in claim 13, wherein the biocomponentis an enzyme.
 15. A sensor as in claim 14, wherein the enzyme is akinase or dehydrogenase, or L(+) lactic dehydrogenase suitable formeasuring pyruvate.
 16. A sensor as in claim 15, wherein the conductingpolymer is a poly(pyrrole).
 17. A sensor as in claim 13, wherein theaffinity component is an antibody, antigen, lectin or receptor.
 18. Amethod for enhancing the sensitivity of a sensor said method comprising:associating an immobilised affinity component capable of interactingwith an analyte species with a conducting polymer such that theinteraction of the affinity component and the analyte induces a changein the electrical properties of the polymer, applying AC signals to thepolymer, and detecting a response of the polymer to the applied signalsto detect the induced change, wherein the impedance of the polymer ismeasured over a wide frequency range, at last one peak in therelationship between frequency and impedance change resulting fromexposure of the polymer to the analyte is detected, and the sensor isoperated at a frequency or frequencies corresponding to a detected peakor peaks.